I. Field of the Invention
The present invention relates to methods of estimating imaged object attenuation maps and more particularly to the completion of truncated attenuation maps by estimating missing portions using Maximum-Likelihood Expectation Maximization (MLEM) reconstructive techniques.
II. Background Information
Nuclear medicine is a unique medical specialty wherein radiation is used to acquire images, which show the function and anatomy of organs, bones or tissues of the body. Radiopharmaceuticals are introduced into the body, either by injection or ingestion, and are attracted to specific organs, bones or tissues of interest. Such radiopharmaceuticals produce gamma photon emissions, which emanate from the body and are captured by a scintillation crystal, with which the photons interact to produce flashes of light or “events.” Events are detected by an array of photo detectors, such as photomultiplier tubes, and their spatial locations or positions are calculated and stored. In this way, an image of the organ or tissue under study is created from detection of the distribution of the radioisotopes in the body.
One particular nuclear medicine imaging technique is known as Positron Emission Tomography, or PET. PET is used to produce images for diagnosing the biochemistry or physiology of a specific organ, tumor or other metabolically active site. Measurement of the tissue concentration of a positron emitting radionuclide is based on coincidence detection of the two gamma photons arising from positron annihilation. When a positron is annihilated by an electron, two 511 keV gamma photons are simultaneously produced and travel in approximately opposite directions. Gamma photons produced by an annihilation event can be detected by a pair of oppositely disposed radiation detectors capable of producing a signal in response to the interaction of the gamma photons with a scintillation crystal. Annihilation events are typically identified by a time coincidence between the detection of the two 511 keV gamma photons in the two oppositely disposed detectors, i.e., the gamma photon emissions are detected virtually simultaneously by each detector. When two oppositely disposed gamma photons each strike an oppositely disposed detector to produce a time coincidence event, they also identify a line of response, or LOR, along which the annihilation event has occurred. An example of a PET method and apparatus is described in U.S. Pat. No. 6,858,847, which patent is incorporated herein by reference in its entirety.
After being sorted into parallel projections, the LORs defined by the coincidence events are used to reconstruct a three-dimensional distribution of the positron-emitting radionuclide within the patient. In two-dimensional PET, each 2D transverse section or “slice” of the radionuclide distribution is reconstructed independently of adjacent sections. In fully three-dimensional PET, the data are sorted into sets of LOR, where each set is parallel to a particular detector angle, and therefore represents a two dimensional parallel projection p(r, s, φ, θ) of the three dimensional radionuclide distribution within the patient—where “r” and “s” correspond to the radial and axial distances, respectively, of the LOR from the center of the projection view and “φ” and “θ” correspond to the azimuthal and polar angles, respectively, of the projection direction with respect to the z axis in (x, y, z) coordinate space (in other words, φ and θ correspond to a particular LOR direction).
Coincidence events are integrated or collected for each LOR and stored in a sinogram. In this format, a single fixed point in the emitter distribution f(x, y) traces a sinusoid in the sinogram. Each row of a sinogram contains the LOR data for a particular azimuthal angle φ; each element of the row corresponds to a distinct radial offset of the LOR from the center of rotation of the projection. Different sinograms may have corresponded to projections of the tracer distribution at different coordinates along the scanner axis and/or different polar angles with respect to the scanner's axis.
FIG. 1 shows an embodiment of an exemplary PET system. A subject 4, for example a patient, is positioned within a detector ring 3 comprising scintillation photon detectors (such as PMT, APD, SiPM . . . ) 5. In front of the scintillation photon detectors 5 are individual crystals 8, also called detectors 8. A group of four scintillation photon detectors may have an array of detectors 8 in front of them. For example, an array of eight by eight or thirteen by thirteen detectors 8 (crystals) is possible, but any other array may be selected. Each detector 8 may be an individual crystal in front of respective scintillation photon detectors. As noted, during an annihilation process two photons 7 are emitted in diametrically opposing directions as schematically illustrated in circle 6. These photons 7 are registered by the PET as they arrive at detectors 8 in the detector ring 3. After the registration, the data, resulting from the photons 7 arriving at the detectors 8, may be forwarded to a processing unit 1 which decides if two registered events are selected as a so-called coincidence event. All coincidences are forwarded to the image processing unit 2 where the final image data may be produced via mathematical image reconstruction methods. The image processing unit 2 may be connected to a display for displaying one or more processed images to a user.
Positron emission tomography provides quantitative images depicting the concentration of the positron emitting substance throughout the patient. The accuracy of this quantitative measurement depends in part on the accuracy of an attenuation correction which accounts for the absorption of some of the gamma rays as they pass through the patient. The attenuation correction factors modify the sinogram, which contains the number of annihilation events at each location within the field of view.
Attenuation is the loss of detection of true coincidence events because of their absorption in the body or due to their scattering out of the detector field of view. Attenuation problems are greater with PET imaging compared to traditional nuclear medicine SPECT imaging. Even though the photons are of greater energy than those used in SPECT imaging, in PET imaging two photons must escape the patient simultaneously to be detected as a true event and the total photon path distance through the object/patient from emission to detection is greater with a PET camera than with a SPECT camera. The loss of true coincidence event detection due to attenuation in PET imaging can range between 50 to 95%, especially great in a larger person.
Loss of counts due to attenuation increases image noise, image artifacts, and image distortion. Without attenuation correction, significant artifacts which may occur on whole-body PET scans include: (1) prominent activity at body surface edges due to relative low attenuation at the surfaces compared to deeper structures, (2) distorted appearance of areas of intense activity (e.g. urinary bladder) due to variable degrees of attenuation in different directions of activity originating from these areas, and (3) diffuse, relatively increased activity in tissues of relatively low attenuation (e.g. lungs). Therefore, attenuation correction of data is necessary for accurate qualitative (i.e. visually normal, increased, or decreased) and quantitative (e.g. standardized uptake values or SUVs for FDG) measurements of radio-tracer activity.
In imaging systems integrating PET and CT imaging modalities, x-rays from a CT scan are used to construct an attenuation map of density differences throughout the body that may then be used to correct for the absorption of the photons emitted from radio-tracer decay. Attenuation is much more likely in the center of the body and therefore non-attenuation-corrected images will show diffusely lower levels of activity deep in the body in comparison to the skin surface. The attenuation correction process essentially “adds counts back” into areas that are more attenuated due to their being deeper or being surrounded by relatively dense structures. Similarly, it essentially “subtracts counts” from areas that are attenuated much less than all other tissues (e.g. lungs and body surfaces). Both attenuation-corrected and non-attenuation-corrected data sets are provided for review and both should be examined by the interpreter. Reviewing both data sets sometimes allows confirmation of an abnormality or confirmation of the benignity of a process which might have been incorrectly assessed based on review of one set alone.
In imaging systems integrating PET and CT imaging modalities, a fundamental hurdle that must be overcome to create an attenuation map is the truncation of the CT image resulting from the CT imaging portion of the system having a smaller field of view than that of the PET camera. This field of view problem also exists in imaging systems that integrate PET and MR imaging modalities, resulting in an even more truncated MR image than that of a CT image resulting from an integrated PET and CT imaging system. The field of view of the PET camera within an integrated or a hybrid imaging system is approximately 60 cm. The field of view of the CT imaging modality of an integrated or a hybrid imaging system is approximately 50 cm. It is possible to overcome the truncation problem in the PET-CT integrated system by collecting image data for a patient that has their arms up over their head during the imaging process which for some patients removes truncation. A patient holding their arms over their head is not a solution for PET-MR systems. The field of view of an MR imaging modality of an integrated or a hybrid PET-MR imaging system is approximately 40-45 cm. In the MR-PET integrated imaging system the conversion of the MR image values to linear attenuation coefficients at 511 keV adds another layer of complexity, because the MR imaging signal does not correlate with electron density.
It is also contemplated that the PET-MR imaging system may be comprised of a separate PET system such as the system disclosed in FIG. 1 and a separate MR system such as that illustrated in FIG. 2. According to FIG. 2, a magnetic resonance system has a base body 11. The base body 11 embodies a magnet system by means of which magnetic fields can be generated in an excitation region 12. The magnet system includes at least one basic magnet 13 for generation of a temporally static basic magnetic field that is spatially at least essentially homogeneous within the excitation region 12. The magnet system furthermore includes a whole-body antenna 14 by means of which a radio-frequency magnetic field can be generated that is at least essentially homogeneous in the entire excitation region 12. The magnet system normally additionally includes gradient magnets for generation of gradient fields and a shielding magnet. The magnetic resonance system includes a patient bed 15 that can be moved in a travel direction z over a travel region relative to the base body 11. The travel region is determined such that—as viewed in the travel direction z—each point of the patient bed 15 can be positioned in the excitation region 12. Since the excitation region 12 normally extends over an excitation region length I which is approximately 40 to 60 cm in the travel direction z and the patient bed exhibits a length L on the order of 2 m, the travel region length of the patient bed 15 is thus inevitably a multiple of the excitation region length I. Due to the mobility of the patient bed 15, an examination subject 16 (normally a person 16) can be brought into the excitation region 12 by corresponding movement of the patient bed 15. It is possible to acquire the emitted magnetic resonance signal by means of the whole-body antenna 14 and to feed it to an evaluation device 17 by which the magnetic resonance signal can be evaluated. However, only a qualitatively low-grade reconstruction of the examination subject 16 is possible in this manner. Local coils 18 by means of which a significantly higher-grade magnetic resonance signal can be acquired (even if only over a small volume per local coil 18) are therefore normally arranged on the examination subject 6.
There is a need for a system and methodology that integrates the use of both PET imaging and an alternative imaging modality, such as CT or MR, into the generation of image attenuation maps that solves the problems caused by the smaller field of view of the alternative imaging modality. Such a system and method needs to provide a solution for the above described attenuation correction problem resulting from a lack of correlation of the alternative imaging modality with the electron density of the emissions data resulting from PET imaging. In a system and method using data sets from both PET and an alternative imaging modality, as the gamma-quanta traverses everything that is inside the system's detector ring before being counted, there is gradual loss of intensity attenuated by the objects within the field of view (FOV). This attenuation must be accounted for through some form of attenuation correction in order to generate clinically relevant image quality. The ability to correct attenuation is not a new concept as it has been successfully accomplished in previous systems also allowing for accurate scatter correction.
The estimation of human attenuation from using MR image data is a difficult task. First, as a preliminary manner, when the alternative imaging modality is MR, the MR image or MR signal has little correlation with the electron densities or the associated linear attenuation coefficients (LAC) of human tissues at the annihilation radiation energy of 511 KeV, the source of radiation in PET. In addition, the transaxial MR field of view is usually a lot smaller (40-45 cm) than the PET field of view (60 cm) for several reasons, e.g. the inhomogeneity of the main field and the non-linearity of the gradients. There is a need for a system and method using data sets from both PET and an alternative imaging modality that incorporates a method of obtaining accurate LAC-values for an image outside the MR image visible field of view that are derived from truncated MR image data based on PET emission data collected for the object being imaged. Currently, MR based estimation of a PET attenuation map is done by segmenting the MR image into different tissue types and assigning the corresponding LAC values, as set forth in US patent Pub. No. US 2008135769. However, this approach alone is not sufficient for estimating the attenuation map outside the MR FOV.
Recently there have been efforts to use maximum-likelihood expectation maximization (MLEM) algorithms to simultaneously reconstruct a complete attenuation map from the PET data emissions data. The success was limited because there was too much missing information. The present invention provides a solution to this problem. This class of algorithms usually converges to a local maximum and therefore has to be constrained to obtain the desired solution. Constraints are generally given by prior information on the smoothness of the emission and attenuation images or allowing discrete values for the attenuation coefficients. Other approaches for MR-based attenuation correction make use of an atlas, model or reference image with known attenuation, e.g. from some co-registered corresponding CT, PET transmission image or body contours derived from optical 3D scanning. The actual MR image is then registered to the atlas or reference with known attenuation and the actual attenuation map is deduced from the registration information.
Notwithstanding the above, the methodology for deductively generating an attenuation map outside the MR imaging modality field of view using an atlas is not particularly straightforward. Accordingly, there is a need for a straightforward method of deducing a complete attenuation map using a truncated image captured by imaging modalities having a field of view that is smaller than that which is necessary to capture a full image.